The present invention relates to systems and methods for emission tomography and, more particularly, to systems and methods for multiple-detection (MD) enhanced emission tomography that provides an increase in the performance of current emission tomography scanners by allowing the counting and effective use for image reconstruction of coincidences involving three or more photons. Such coincidences include, but are not limited to, coincidences caused by inter-detector scattered photons, random coincidences involving more than two photons, and multiple-photon events caused by radionuclides that emit prompt gamma rays in coincidence with positron emission.
There are a variety of emission tomography imaging systems and methods. One clinically important example is positron emission tomography (PET) which, generally, utilizes an administered radionuclide to acquire two-dimensional and three-dimensional tomographic images of a target area or organ of interest in a subject. More specifically, such radionuclides are employed as radioactive tracers called “radiopharmaceuticals” by incorporating them into substances, such as glucose or carbon dioxide, or molecules specifically designed to bind a predetermined target (for example, an antibody targeting a cell surface protein). These radiopharmaceuticals are then administered to the patient where they become involved in biological processes such as blood flow; fatty acid and glucose metabolism; and protein synthesis. Through a respective biological process, the radiopharmaceuticals accumulate in, or otherwise target, the area or organ of interest in the subject. By measuring or identifying photons emitted from the area or organ of interest by the accumulated or targeted radiopharmaceutical, clinically useful biological and physiological information can be acquired from the area or organ of interest.
For example, in PET, as the injected radioactive tracer decays, it emits positrons. The positrons travel a very short distance before they encounter an electron and, when this occurs, the positrons are annihilated and converted into two high-energy photons, or gamma rays. This annihilation event is characterized by two features that are pertinent to PET imaging. Namely, each gamma ray has an energy of 511 keV and the two gamma rays are directed in substantially opposite directions. An image is created by determining the number of such annihilation events at each location within the scanner's field of view (FOV).
To create such an image, typical PET scanners consist of one or more rings of detectors which are positioned to encircle the patient. Coincidence detection circuits connected to the detectors record only those photons that are detected simultaneously by two detectors located on opposite sides of the patient and that fall within an energy acceptance window around 511 keV. The number of such simultaneous events indicates the number of positron annihilations that occurred along a line joining the two opposing detectors. Within a few minutes, millions of events can be recorded to indicate the number of annihilations along lines joining pairs of detectors in the ring. These numbers are employed to reconstruct an image using well-known tomographic reconstruction techniques.
For example, current clinical (and most preclinical) PET scanners and systems usually consist of a ring 100 of block detectors 102 for detecting emitted photons, typically in circular, such as the array shown in FIG. 1, or in hexagonal or octagonal arrays. Block detectors 102 typically include a piece of scintillator material that converts the energy deposited by gamma rays into visible light. The scintillator material is usually segmented into many scintillation crystal elements configured in an array, which is read out by a number of individual photo-detectors (typically, photo-multiplier tubes (PMTs), a position-sensitive photo-multiplier tube (PS-PMT), or silicon photo-multipliers (Si-PM)) that convert the light emitted by the scintillation material into electrical signals whose magnitude is proportional to the energy deposited by the gamma rays in the scintillator material. By combining the output signals of the photon detector(s) of the block detector, it is possible to determine the single crystal in which the detected photon interacted and the energy deposited by such photon.
Although block detectors have been demonstrated as the most cost-effective solution for the implementation of PET scanners, these detectors also present some drawbacks. One drawback is that, since each detector element is a block, if several photons interact simultaneously on the same block and the added energy of those photons is within a predefined energy acceptance window (around 511 keV), it is not possible to determine from the output signals of the detector if they were produced by the interaction of a single photon (thereby presenting useful information) or by the interaction of multiple photons (thereby presenting distorted or non-useful information).
In addition, as shown in FIG. 1, the ring of block detectors 100 of a PET scanner includes individual detectors that are operated in coincidence with a fan beam 104 of block detectors on the opposite side of the ring 100. The inner circle 106 formed by edges of all such fan beams defines the useful field of view. Data is usually recorded simultaneously for all possible fan beams, and the PET scanner will produce an output whenever two photons are detected in opposite block detectors of a fan beam 104 within a specified coincidence timing window (for example, in the range of hundreds of picoseconds to tens of nanoseconds) and when both events fall into a predetermined energy window (usually from 511 keV−ΔE1 to 511 keV+ΔE2 where ΔE1 and ΔE2 are chosen as function of the energy resolution of the block detectors). Any such events are called prompt coincidences, but can be of three specific types: true coincidences, scatter coincidences, and random coincidences. In some cases, prompt coincidences (that is, any coincidence involving two photons) may be simply referred to as coincidences, in comparison to multiple-detection (MD) coincidences or multiple interaction photon (MIP) coincidences, which include coincidence events involving three or more detected photons, as further described below.
True coincidences occur when two photons 200 and 202 produced from the same positron annihilation 204 are detected within the time and energy windows of the system, as shown in FIG. 2A. Scatter coincidences occur when at least one of the photons undergoes scattering in the object under study, for example, where the photon loses a fraction of its total energy in the scatter interaction with the object before its detection. The scatter coincidence is thus detected in a pair of detectors that are non-collinear with the originating annihilation, as shown in FIG. 2B. Random coincidences, also known as accidental coincidences, occur when annihilation photons 200a and 202b from two unrelated positron annihilation events 204a and 204b are detected in opposite detectors, as shown in FIG. 2C. True coincidences produce valid information, while both scatter coincidences and random coincidences produce distorted information. In particular, scatter and random coincidences yield incorrect positional information, as shown by the dotted lines in FIGS. 2B and 2C, and contribute to a relatively uniform background noise in the resulting image, which results in a loss of contrast.
With respect to scatter coincidences, such events are typically assumed to occur only due to scattering within the patient, as shown in FIG. 2B, and current PET systems include scatter correction procedures based on this assumption. However, there are a large number of events in which Compton scattering occurs in the block detectors of the scanner, as shown in FIGS. 3A and 3B, depositing a fraction of the total energy of the photon in each interaction. In particular, FIG. 3A illustrates a scatter event where one of the photons from an annihilation event (photon A) interacts by photoelectric effect depositing energy in a detector within the acceptance energy window of the scanner, and the other photon (photon B) interacts by Compton scatter in another detector, where it deposits some of its energy, with the scattered photon (photon C) escaping from the detector ring. Consequently, the scanner will process such an event as a prompt coincidence and will accept or discard the event depending on the energy of photon B. More specifically, if the energy of photon B is within the scanner's energy acceptance window, it will be labeled as a true coincidence event and accepted. If the energy of photon B is not within the energy acceptance window, it will be labeled as a scatter coincidence event and discarded.
FIG. 3B illustrates an inter-detector scatter (IDS) event, which is a specific case of an MD event (that is, an event involving more than two detected photons). The IDS event of FIG. 3B occurs when one of the photons from an annihilation event (photon A) interacts by photoelectric effect depositing energy in a detector within the acceptance energy window of the scanner (that is, 511 keV−ΔE1 to 511 keV+ΔE2), and the other photon (photon B) interacts by Compton scattering in another detector. Photon B deposits some of its energy in the detector it is incident upon, and the scattered photon (photon C) produced by the Compton scattering event deposits energy in another detector. Consequently, multiple photons (three, in this example) are detected within the time window of the scanner and this MD event could be processed to obtain useful information.
With respect to random coincidences, events (that is, random MD events) can involve more than two photons from at least two different decays within accepted energy and timing windows. When random MD events are detected, current PET scanners either reject the multiple detected photons or erroneously select one or more photon pairs and respective lines of response (for example, as a function of the timing and/or energy of the detected photons). When the photons from the MD event originate from the same annihilation process, the MD event can be processed to obtain useful information. For example, FIG. 3C illustrates a random MD event where three photons, photon A, photon B, and photon C, are detected within the coincidence and energy windows of the scanner. In this example, photons A and B come from the same annihilation event, whereas photon C originates from a different annihilation event. As photons A and B arise from the same positron-electron annihilation, the line A-B contains useful information while lines A-C and B-C do not. Alternatively, some detected random MD events originate from N different annihilations (where N is the number of photons involved in the MD event). While such an event has a lower probably of occurring, the processed event would not provide any useful information. For example, FIG. 3D illustrates a random MD interaction in which three photons, photon A, photon B, and photon C, are generated by three different annihilations. In this example, none of the possible lines of response (that is, line A-B, line A-C, or line C-B) provide useful information.
In addition, there are several radionuclides of interest to emission tomography that emit prompt gamma rays in coincidence with the emission of a positron. More specifically, a radionuclide decays by positron emission and, after a short delay (in the range of picoseconds), one or more prompt gamma rays are also emitted. This results in MD events (considered positron-gamma MD events) involving simultaneous detection of more than two gamma rays coming from the same nuclear decay. Examples of such radionuclides that are capable of causing such events (considered positron-gamma emitters) include, but are not limited to, iodine-124 (124I), bromine-76 (76Br), yttrium-86 (86Y), rubidium-82 (82Rb), and technetium-94m (94mTc). FIG. 3E illustrates an example positron-gamma MD event that may occur when using a positron-gamma emitter with a state-of-the-art PET scanner. In this example, photon A, photon B, and prompt gamma ray C are generated from an annihilation. If the energy of prompt gamma ray C is within the energy acceptance window of the scanner, the event may be processed similar to the random MD events described above with respect to FIGS. 3C and 3D, where a line of response is selected as a function of the timing resolution and/or energy resolution of the scanner. If the energy of prompt gamma ray C is over or under the energy acceptance window of the scanner, the scanner may treat the event as a true coincidence and select the appropriate line of response (that is, line A-B) using suitable criteria.
In current clinical and preclinical PET scanners, no viable information is obtained when the above-described MD events occur (that is, the IDS event of FIG. 3B, the random MD events of FIGS. 3C and 3D, and/or the positron-gamma MD event of FIG. 3E) because multiple detections are not identified as valid events by the scanner's coincidence system and, thus, are rejected or erroneously included as double-coincidence events. Although, in some cases, the IDS event shown in FIG. 3A may be detected and processed in the same fashion as scatter events that have undergone scattering in the object, such as that shown in FIG. 2B. In other words, data collected for events comprising more than two detections is usually thrown out and only data from prompt coincidences (including true coincidences, in-body scatter coincidences, random coincidences, crystal scatter coincidences with two detection events, and in general any type of two-photon coincidence events within time and energy acceptance windows) are used to compose images, thus limiting the potential sensitivity of the system and quality of the resulting images.
Therefore, it would be desirable to have a system and method for emission tomography imaging that controls for data collected from MD events and, additionally, may effectively use these events for generating improved images without introducing artifacts.